Optical coherence tomography for measurement on the retina

ABSTRACT

An optical coherence tomograph that provides wavelength tunable source radiation and an illumination and measurement beam path, a dividing element that divides source radiation into illumination radiation and reference radiation, and collects measurement radiation. The illumination and measurement beam path has scanner. A detection beam path receives measurement radiation and reference radiation and conducts them onto at least one flat panel detector in a superposed manner. A beam splitter separates the measurement radiation from the illumination radiation. The beam splitter conducts the separated measurement radiation to the detection beam path and sets the numerical aperture of the illumination of the illumination field in the eye. An optical element sets the numerical aperture with which the measurement radiation is collected in the eye and a multi-perforated aperture defines the size of an object field and a number of object spots, from which the measurement radiation reaches the flat panel detector.

RELATED APPLICATIONS

The present application is a National Phase entry of PCT Application No.PCT/EP2015/051834, filed Jan. 28, 2016, which claims priority from DEPatent Application No. 10 2015 101 251.0, filed Jan. 28, 2015, both ofsaid applications being hereby incorporated by reference herein in theirentirety.

FIELD OF INVENTION

The invention relates to an optical coherence tomograph for examining asample, in particular an eye.

Further, the invention relates to a method for optical coherencetomography for examining a sample, in particular an eye.

BACKGROUND

In optical coherence tomography OCT systems the lateral resolution (xand y) is defined by the numeral aperture (NA) of the optical systemused. The axial resolution (z), however, is calculated from aninterference pattern and as a rule is much greater than the depth offield of the imaging, which in turn depends on the numerical aperture,more precisely is proportional to 1/NA². In the usually used Fourierdomain OCT, which uses a broadband or wavelength-adjustable radiationsource, the depth resolution is inversely proportional to the spectralbandwidth, more precisely proportional to λ²/Δλ, wherein λ is theaverage wavelength and Δλ is the bandwidth. Optical coherence tomography(OCT) is an established method for imaging the eye in ophthalmology. Itmakes a three-dimensional imaging possible, which is very useful for thediagnosis of eye diseases and the progression thereof. To be named inparticular here are diseases of the retina, such as glaucoma orage-related macular degeneration.

To measure the retina of the human eye both a high lateral resolutionand a high axial resolution are needed. At the same time the detectableand thus illuminated volume is to be as large as possible in terms ofdepth (along the optical axis); this requires a small numerical aperture(NA) of the optical system. The lateral resolution requires a largenumerical aperture. Thus, ultimately, in the state of the art, theextent of the area accessible in terms of depth and the lateralresolution are linked to each other via the numerical aperture of theoptical system and cannot be set independently of each other.

An OCT based imaging method is known from US 2014/0028974 A1. A line isprojected through an imaging system onto a sample. The backscatteredradiation is combined in an interfering manner with reference radiationand guided to a detector, wherein a confocal filtering is carried out inone direction. An astigmatic optical system is used for this. The depthresolution is obtained by optical coherence tomography. In the case of aspectroscopic analysis of the radiation, a two-dimensional detector isused one dimension of which is used for the confocal filtering withrespect to the illuminated line area and the other dimension of whichresolves the spectral information. Lateral resolution and accessibledepth area are also linked in the approach according to US 2014/0028974A1.

WO 2014/0179465 A1 describes an OCT which operates in the spectraldomain, thus analyses the interference of radiations with aspectrometer. The light source emits a bundle of light which consists ofa plurality of parallel individual bundles which are imaged onto thesample through the objective lens. A reference arm also guides severalparallel individual bundles, with the result that in the end eachindividual bundle is guided through the device according to the OCTmeasurement principle and also analysed on the spectrometer. This deviceis very complicated to align.

In a scanning OCT system the accessible diameter of the pupil of the eyeis usually between 1 mm and 1.5 mm. This results in a lateral resolutionof approximately 15 μm and an area detectable in terms of depth with anextent of 3 mm. A better lateral resolution would be achieved with ahigher numerical aperture of the optical system. However, at the sametime, the depth-detectable area would thus be reduced. In addition,aberrations increase with the numerical aperture. In known OCT systemswhich use a diameter of up to 1.5 mm in the pupil of the eye astigmatismand coma increase for larger pupils even if the defocusing isdisregarded as a higher-order aberration. A diffraction-limitedresolution can therefore not be achieved.

For particular applications, in particular for the diagnosis ofage-related macular degeneration, a high lateral resolution is desired.In order to diagnose an early stage of this disease, a lateralresolution of approximately 5 μm is needed. At the same time, a depthmeasurement area that can be sampled of approximately 3 mm is required,as it is assumed that age-related macular degeneration is accompanied byblood vessel growth in deeper layers of tissue. In order to detect suchvessels, a good signal-to-noise ratio is additionally needed.

SUMMARY OF THE INVENTION

The object of the invention is to specify an optical coherence tomographfor measurement on the retina of the human eye, in which the lateralresolution is improved without limiting the accessible depth area at thesame time.

The invention combines several features in order to obtain, byapplication of optical coherence tomography, a three-dimensional imagewhich has a particularly good resolution laterally, i.e. transverse tothe optical axis, and at the same time can cover a very large depth areaaxially, i.e. along the optical axis, without the need to adjustfocusing elements or lenses during the measurement process.

The invention provides for a multi-spot holoscopy OCT. One aspect of theinvention comprises the sample being illuminated and imaged at aplurality of object spots simultaneously, wherein the imaging is done inparallel by filtering the radiation by use of a multi-hole diaphragmduring the detection. Each object spot is imaged onto a dedicateddetector pixel area. In embodiments an oversampling is done, i.e. theresolution of the detector pixel area of each object spot is greaterthan the diffraction limit of the optical imaging actually allows. Inexample embodiments an image correction is then obtained from theresolved intensity distribution.

One embodiment of the coherence tomograph comprises: an illuminationdevice for providing source radiation the wavelength of which istunable, an illumination and measurement beam path which has a dividingelement for dividing the source radiation into illumination radiationand reference radiation, illuminates an illumination field in the eyewith the illumination radiation and collects illumination radiationbackscattered in the eye as measurement radiation, wherein theillumination and measurement beam path comprise a scanner for adjustingthe lateral position of the illumination field in the eye and a frontoptical system, a reference beam path which provides an optical pathlength, which corresponds to an optical path length from the splittingelement to the illumination field and back, for the reference radiation,a detection beam path which receives the measurement radiation from theillumination and measurement beam path and the reference radiation fromthe reference beam path and guides them, superimposed, onto an areadetector.

In one embodiment, the optical coherence tomograph comprises: anillumination device for providing source radiation wherein theillumination device is tunable regarding the wavelength of the sourceradiation, an illumination and measurement beam path which comprises adividing element for dividing the source radiation into illuminationradiation and reference radiation, projects the illumination radiationto an illuminated field in the eye and collects illumination radiationbackscattered in the eye as measurement radiation from an object field,wherein the illumination and measurement beam path comprises a scannerfor adjusting the lateral position of the illumination field and theobject field in the eye and a front optics. A reference beam path whichprovides an optical path length for the reference radiation which pathlength corresponds to an optical distance from the splitting element tothe illumination field and back, a detection beam path receiving themeasurement radiation from the illumination and measurement beam pathand the reference radiation from the reference beam path andsuperimposes them at a point of superimposition and guides them to a 2Ddetector.

In embodiments the illumination and measurement beam path furthercomprises a beam splitter for splitting off the measurement radiationcollected from the eye from the illumination radiation projected to theeye, wherein the beam splitter guides the split off measurementradiation to a detection beam path, and a light-distributing elementwhich distributes the illumination radiation into spots, in order toilluminate the illuminated field with a multi-spot pattern.

Some embodiments have the detection beam path further comprising: anintermediate image plane, an optical element which only acts on themeasurement radiation and which cooperates with the front optics andsets the numerical aperture with which measurement radiation iscollected from the eye, and a diaphragm which is upstream of the 2Ddetector and is arranged in or close to the intermediate image plane anddefines the size of the object area, wherein the diaphragm upstream ofthe detector is formed as a first multi-hole diaphragm and a firstmulti-lens array, which bundles the radiation emerging from each hole ofthe first multi-hole diaphragm onto a dedicated pixel area of thedetector for each spot which pixel areas each have a spatial resolutionof, for example, 4 to 100 pixels in one direction, in another example,an 2D area of 5 to 50 pixels or of 5 to 40 pixels per direction, isarranged between this multi-hole diaphragm and the 2D detector.

The method comprises providing source radiation, sweeping the wavelengththereof and dividing the source radiation into illumination radiationand reference radiation, projecting the illumination radiation to anilluminated field in the eye and collecting illumination radiationbackscattered in the eye as measurement radiation from an object fieldin the eye, wherein a scanner for adjusting the lateral position of theilluminated field and of the object area in the eye and front optics areused, delaying reference radiation in a reference beam path, andsuperimposing the measurement radiation with the reference radiation anddetecting an interference signal of the superimposed radiations with an2D detector.

Some embodiments comprise splitting the measurement radiation collectedfrom the eye from the illumination radiation projected to the eye,projecting illumination and collecting the measurement radiation atindependent numerical apertures by using an optical to element whichonly acts on the illumination radiation and which cooperates with thefront optics and by using an optical element which only acts on themeasurement radiation and which cooperates with the front optics.

The radiation of the imaged object spot is superimposed with referenceradiation according to the OCT principle, with the result that thedetector spots receive an interference signal between measurementradiation and reference radiation.

A 2D detector is used which samples the object spots on the retina. Themulti-hole diaphragm of the optical imaging defines these object spots,and the 2D detector is matched, as a whole, to the size of the areacovered by the diaphragm with holes and, with respect to its spatialresolution or pixels, to the size of the holes.

A sweepable light sources used, and detector without any dispersiveelement takes the place of a spectrometer. The detector resolves eachsingle spot by several pixels in the spatial domain x, y or thefrequency domain kx, ky. In contrast to WO 2014/149465 A1, the frequencydomain is filled in kz in that several frames are taken during one sweepperiod of the light source. By detecting each spot in x, y or kx, ky, ahigher numerical aperture (NA) can be utilized because wavefrontcorrections can be done numerically and even after the frames had beentaken, resulting in higher resolution and larger depth area.

In embodiments, the illumination beam path and the detection beam pathare the same and share, in particular, a front optic. Even then the beampath is configured such that the illumination by the illuminationradiation and the collection of the backscattered measurement radiationhave independent, e.g. different, numerical apertures. Thus theillumination is provided with a numerical aperture which illuminates anaxially large area, with the result that the collected measurementradiation originates from a comparatively large depth area andconsequently the OCT principle obtains an image over a large depth area.The numerical aperture of the collection of the measurement radiation,thus of the imaging of an object area, is set independently of thenumerical aperture of the illumination, e.g. may be larger. A highlateral resolution is thereby combined with a large illuminated deptharea.

The detector is a two-dimensional detector comprising pixels. Inembodiments, the pixel number lies between 4 and 100 pixels perdirection and per hole of the multi-hole diaphragm, in another examplebetween 5 pixels and 40 pixels. Such pixel numbers proved to beadvantageous for sampling each object spot both with respect toresolution and with respect to signal-to-noise ratio and possible imageerror corrections.

The aberrations that the eye generates are particularly important forimage error correction. Since the invention decouples the numericalapertures of illumination and detection, some embodiments carry out thedetection, i.e. the imaging of the object area on the retina, atnumerical apertures, at which aberrations of the eye might become aproblem. The spatial resolution at which the 2D detector detects eachobject spot makes it possible, as explained below, to correct theaberrations in particular if the detector is arranged in a conjugatedpupil plane of the imaging beam path. If the detector does not lie in apupil plane, an aberration correction is likewise possible if thedetected signal is converted to relate to a pupil plane, as is known forholograms in the state of the art.

In the object plane and the (intermediate) image planes of a beam paththe image information is pure location information. Imaged structurescan be noted in (intermediate) image planes. There they take the form ofintensity differences. In pupil planes the image information is pureangle information. Here, the angles of the incident beams encode theimage information. This has the known effect that a cross-sectionalalteration in a pupil influences exclusively the image brightness, butnot the image size. For this reason the human eye iris lies in the pupilplane, with the result that the human eye adapts with respect tobrightness by narrowing or widening of the iris. When the plane of thepupil of the eye is discussed in this description, the iris plane ismeant. An imaging beam path images an object from the object plane ontoan image in the image plane (e.g. the location of a detector). Becauseof the laws of imaging, there always exists a pupil between for examplethe object plane and an intermediate image plane. Similarly, therealways exists an intermediate image plane between two pupil planes.Likewise, in this description, planes which are located between theplane of the pupil of the eye and the detector are called conjugatedpupil planes, as they are conjugated to the plane of the pupil of theeye, predetermined by the optically imaging elements. Where the retinais named as the object here, that is not intended to limit theinvention. Other structures of the eye can equally be imaged as object.

The features of the optical coherence tomography described herein can beused alone or in different combinations for various embodiments. Wherethe following embodiment examples describe particular combinations offeatures, the invention is not limited to such combinations.

The invention achieves a combination of the advantages of a confocalscanning system with the advantages of a spatially sampling detector.The confocal principle of a scanning system supresses scatteredradiation very effectively, whereby an improved signal-to-noise ratio isachieved. At the same time the lateral resolution can be increased byenlarging the aperture on the eye. The invention provides that thenumerical aperture of the illumination is decoupled from the numericalaperture of the detection. This achieves a high lateral resolutionwithout impairing the detectable depth area. The design aimscontradictions in the state of the art (high lateral resolution requireshigh NA, large detectable depth area requires small NA) are thus bothachieved.

Some embodiments of the invention use a confocal multi-hole diaphragm.In this description the term “confocal” not only relates to a diaphragmwhich lies exactly in an (intermediate image) plane conjugated to theobject plane, but also encompasses an arrangement of the diaphragm whichlies within a certain margin of error in front of or behind anintermediate image plane. If the confocal diaphragm does not lie exactlyin the intermediate image plane, but lies close to the intermediateimage plane, then, although a scattered light suppression may possiblybe reduced, the function as confocal diaphragm which defines the objectfield from which the measurement radiation is collected is stillfulfilled. The diaphragm is in or close to an intermediate image planeas long as it is in a distance from the intermediate image plane notmore than three times the imaging depth; a distance of not more than onetimes the imaging depth is for example, preferred. The imaging depth isalso called “depth of focus” in the literature and defines an axial areain the image space, i.e. on the intermediate image plane of an opticalsystem in which a sufficiently sharp image forms in a detection plane.Diffraction spots are detected as a dot in the area of the depth offocus. The area in the object space conjugated to the depth of focus isthe depth of field. The depth of field is a measure of the extent of thesharp area in the object space and is given by lambda/(NAo)², whereinNAo denotes the numerical aperture in the object space. The depth offocus on the intermediate image plane results, analogously to the depthof field, from the numerical aperture by lambda/(NAz)²; NAz is thenumerical aperture in the intermediate image plane, which is calculatede.g. from NAo by of the imaging scale. In the above calculations, themaximum wavelength of the measurement radiation can be used aswavelength in the intermediate image plane.

In ophthalmological applications, the invention has the advantage thathigher light intensities can be radiated into the human eye, becausethey are distributed onto larger cross sections of the anterior chamberof the eye. To obtain that advantage the illumination beam pathcomprises structure which illuminates the retina with a multi-spotpattern and places the pupil of the illumination within the eye and, insome example embodiments, illuminates the full diameter of the iris(approx. 4.2 mm) in the eye. Some example embodiments place theeffective position of the illumination pupil in the plane of the iris.This can be done by using a field lens, alternatively also by acorresponding layout of the optical system between the eye and by amulti-spot diaphragm of the illumination beam path.

A pupil position outside the eye would have the result that the bundleof illumination beams does not irradiate through the iris on the opticalaxis for all spots. A small angle between illumination and detection,which can lead to vignetting effects over the measurement depth of theretina, thereby results. These effects can be disadvantageous for highlyprecise applications.

The invention generates several illumination spots on the object, e.g.the retina, from a homogeneous plane illumination wavefront withoutlosses and at the same time utilizes all pixels on the 2D detector.

Various principles come into consideration for the detection. Someembodiments use a single detector or a balanced detection or an off-axisdetection.

It is understood that the features mentioned above and those yet to beexplained in the following are applicable not only in the statedcombinations, but also in other combinations or singly, withoutdeparting from the scope of the present invention.

BRIEF DESCRIPTION OF THE DRAWINGS

The invention is explained in even more detail below by way of examplewith reference to the attached drawings, which also disclose featuresessential to the invention. There are shown in:

FIG. 1 a schematic representation of an optical coherence tomograph(OCT) in a first embodiment,

FIG. 2 a schematic representation of an OCT in a second embodiment,

FIGS. 3a-d schematic representations of different variants of thedetection beam path of the OCT of the first and second embodiments,

FIG. 4 representations to illustrate an aberration correction which canbe used in an OCT of FIG. 1 or of FIG. 2,

FIG. 5 a top view of a detector which can be used in an OCT of FIG. 1 orof FIG. 2,

FIG. 6 a representation to illustrate a depth correction which can beapplied in an OCT of FIG. 1 or of FIG. 2,

FIG. 7 signal intensities of different channels of a detector of an OCTaccording to one of FIG. 1 or 2,

FIG. 8 a representation similar to that of FIG. 7,

FIG. 9 a schematic representation to illustrate the scanning principlein an OCT according to one of FIG. 1 or 2, and

FIGS. 10 and 11 schematic representations similar to that of FIG. 9 toillustrate the generation of a three-dimensional image.

DETAILED DESCRIPTION

FIG. 1 shows an OCT 1 which captures three-dimensional images of aretina 2 of an eye 3. Source radiation of a radiation source 4 which istunable with respect to its wavelength, for example of a correspondinglaser, is coupled into a fiber 5. The source radiation is for example ofinfrared wavelength. In the following description this wavelength rangeis also called “light”. All radiation of the electromagnetic spectrumwhich satisfies the optical laws is to be subsumed under this term.

The fiber 5 feeds a splitter 6, which splits the source radiation into ameasurement arm 7 and a reference arm 8. In the measurement arm 7 afiber 9 follows the splitter 6, and the illumination radiation Bemerging at the fiber end is guided to a beam splitter 11 by means wayof an illumination optical system 10. From there it reaches a frontoptics 12, which bundles the illumination radiation B into a focus whichlies on the retina 2 of the eye 3. The illumination optical system 10and the front optics 12 set, among other things, numerical aperture NAwith which the eye 3 is illuminated. A scanner 13 which deflects thefocus on the retina 2 biaxial and perpendicular to the direction ofincidence, i.e. lateral, is located between the beam splitter 11 and thefront optics 12. The directions of this deflection may be denoted x andy in the following. A z position of the focus can be set by adjustmentof the front optics 12. This is indicated schematically by a doublearrow in FIG. 1.

The illumination radiation in the illumination focus on the retina 2 isbackscattered from different depths within a depth of field range. Thedepth of field range is defined by numerical aperture NA, which isdetermined by the front optics 12 and the illumination optical system 10as well as the optical properties of the eye 3.

Backscattered radiation is collected by the front optics 12 asmeasurement radiation M. To distinguish between the incidentillumination radiation B and the backscattered measurement radiation Mcollected by the front optics 12, these are entered differently inFIG. 1. In the figure the illumination radiation is shown withcontinuous lines, the measurement radiation M with dotted lines. Themeasurement radiation collected by the front optics 12 is guided to thescanner 13. It is descanned here, with the result that after the scanner13 the measurement radiation M is a stationary beam.

Collection of the measurement radiation M is, in fact, imaging of theretina 2. The beam splitter 11 separates the measurement radiation Mfrom the illumination radiation B and guides the isolated measurementradiation 14 to a detector device 17. The detector device 17 will beexplained in more detail later with reference to FIG. 3. It has, amongother things, an optical system which, together with the front optics 12and the optical properties of the eye 3, defines numerical aperture NAof the imaging of the retina 2. In this way, illumination and detectionhave independent numerical apertures. Numerical aperture of theillumination is defined by the illumination optical system 10 and thefront optics 12, numerical aperture of the detection is defined by thefront optics 12 and the detector device 17.

Reference radiation R from the reference arm 8 is also coupled intowards the detector device 17. The reference arm comprises a fiber 20after the splitter 6. In the embodiment shown in FIG. 1 the referencearm 8 comprises a pathlength adjusting device 21, which serves to setthe optical length of the reference arm 8 suitably relative to theposition of the retina 2 of the eye 3. In an embodiment of thepathlength adjusting device 21 couples out the radiation from the fiber20 and guides the radiation through a retroreflector 22, the position ofwhich can be adjusted, as indicated by the double arrow in FIG. 1. Afurther deflecting mirror 23 and optical systems 24, 25 guide thereference radiation R to the detector device 17, which guides thereference radiation R superimposed with the measurement radiation M to a2D detector 19.

The pathlength adjusting device 21 is provided as a free beam path inFIG. 1. This is optional, as is the use of a retroreflector 22. In thestate of the art various measures are known for adjusting the opticalpathlength of a beam.

The interference between reference radiation R and measurement radiationM is implemented to generate an image by optical coherence tomography.As the wavelength of the source radiation is tuned, the Fourier domainprinciple is used for OCT image generation, which is known to personsskilled in the art.

For image generation, OCT 1 comprises a control device C which receivesa signal about the wavelength tuning and the measurement signals of thedetector 19. Optionally, the control device C controls the wavelengthtuning of the radiation source 4 and, therefore, knows the wavelengthcurrently propagating in the system and can thus process the measurementsignals accordingly. The detector 19 receives measurement radiation Mfrom an object field on the retina 2, which field is defined by adiaphragm in the detector device 17 (see FIG. 3). The detector 19comprises pixels and is matched in terms of its size to this diaphragmand samples the intensity distribution in a spatially resolved mannerwith the pixels. In some embodiments the detector 19 lies in an imageplane, i.e. in a plane which is, in the imaging carried out by the frontoptics 12, detector optical system 14 and the further optical elementslocated in-between, conjugated to the plane of the retina 2. Then, theindividual pixels directly provide the location information in theobject field. In other embodiments the detector 19 lies in a conjugatedpupil plane which is conjugated to the plane in which the pupil P of theeye 3 lies. Then, the pixels detect the intensity distribution in thepupil plane and thus the angle information. This can also be used forimage reconstruction, as will be explained below.

For the invention it is important that the scanner 13 shifts the objectfield in the retina 2 and acts not only on the illumination radiation B,but also on the collection of the measurement beams M. A partial imageof the retina thus forms at each position of the scanner 13. Thesepartial images are, as will be explained below, combined to form a totalimage which has a much higher resolution than those known from widefieldOCT.

In the embodiment of FIG. 1 the detector device 17 combines themeasurement radiation M from the measurement arm 7 and the referenceradiation R from the reference arm 8. The detector 19 detects thepattern of interference between measurement radiation M and referenceradiation R. To generate such an interference, in particular theproperties of the radiation source 4 and the path length adaptation, areknown in the art of optical coherence tomography.

The complex amplitudes of the measurement radiation and of the referenceradiation can be written as:

U _(sample) =u _(s) *e ^(iφ) ^(s) and

U _(reference) =u _(r) *e ^(iφ) ^(r) ,

if u_(s) and u_(r) denote the amplitudes and φ_(s) and φ_(r) denote thephases of the signals in the two arms (the subscripts “sample” and “s”refer to the measurement arm, the subscripts “reference” and “r” referto the reference arm).

The detector detects a signal I₁ and, in the case of a “balanceddetection”, which will be discussed later, also a signal I₂:

I ₁ =|U _(sample) +U _(reference)|² =|U _(sample)|² +|U_(reference)|²+2Re{U _(sample) *Ū _(reference)}

|I ₂ =|U _(sample) +U _(reference) *e ^(iπ)|² =|U _(sample)|² =|U_(reference)|²+2Re{U _(sample)*Ureference*e−iπ.

The amplitude of the interference signal is modulated to common-modeportions |U_(sample)|² and |U_(reference)|² is filtered out bycorresponding data analysis or a balanced detection or an off-axisdetection.

FIG. 2 shows a further embodiment for the OCT 1, in which the pathlengthadjusting device is arranged not in the reference arm 8, but in themeasurement arm 7. Pathlength adjusting device 29 is located after thefiber 9 and the illumination optical system 10, again purely by way ofexample by use of a movable retroreflector 30. The embodiment of FIG. 2shows that it makes no difference whether the pathlength adjustingdevice lies in the reference arm 8 or in the measurement arm 7. It isalso possible to provide a pathlength adjusting device in both arms. Theonly important thing is that the interference state between thereference radiation R from the reference arm 8 and the measurementradiation M can be matched to the current measurement task, i.e. theactual position of the object to be measured, in the embodiment examplesdescribed herein the retina 2 of the eye 3.

In FIG. 2 an embodiment is shown in which the front optics 12 is formedin two parts by two imaging elements 12 a and 12 b. This is optional.

For applications at the eye 3 specifications for a maximum allowableillumination intensity on the cornea are to be obyed. If the illuminatedfield is enlarged, more illumination radiation energy can be coupled inonto the eye 3, without exceeding a threshold for the illuminationintensity density. In ophthalmological applications and in the infraredwavelength of usual OCTs, a maximum luminance of approximately 1 mW/mm²must not be exceeded in the anterior chamber of the eye. If an eye pupildiameter of 4.5 mm is illuminated homogeneously, a total ofapproximately 16 mW would be allowable, thus. In order not to allow thedepth-scannable area to become too small, however, the whole pupil P ofthe eye 3 is not utilized for the illumination. Instead, an NA ofapproximately 0.035 (or a pupil diameter of 1.2 mm) is for example usedas upper limit for proper depth detection.

For the tissue of the retina the maximum allowable power is at 1.5 mWfor spots smaller than 1.5 mrad and for a wavelength of 1060 nm. Thishas the result that the 16 mW allowable with respect to the pupil haveto be distributed over an angle of 15 mrad in at least one direction, inorder not to exceed the maximum value of the retina 2. Then the wholesignal intensity would have been maximized, but at the expense of theimage contrast, because scattered light is to be expected for such highintensity radiation under normal widefield illumination conditions.

OCT 1 resolves this conflict of aims by illuminating and detecting theretina simultaneously at several spots spaced apart from each other. Theproblem of scattered light is minimized by the spacing of the spots.FIGS. 3a-d schematically show the elements of OCT 1 which are relevantfor this, wherein all components which relate to the generation andcoupling-in of the reference radiation R have been omitted for the sakeof simplicity. Elements of FIGS. 3a-d which correspond to elements whichhave already been explained with reference to the previous figures areprovided with the same reference signs, with the result that repetitionof their description can be dispensed with. In the representation ofFIG. 3a the coupling-in of the reference radiation R takes place at thesite of the dotted double line. FIGS. 3b-d show different variants forimplementing the coupling-in and also the detection.

Illumination and detection are done in accordance with a multi-spotprinciple in the OCT of FIGS. 1 and 2, as FIGS. 3a-d show. However, thisdoes not apply to the reference beam path 8, which does not guideradiation in several spots. The guiding of the reference radiation isnot shown in FIG. 3. The reference radiation is coupled in at the siteof the dotted double line, thus, in imaging direction of FIG. 3a , afterthe beam splitter 11 and before the optical system 14. This coupling-inis not shown in FIG. 3a for the sake of simplicity. Optical system 14 isoptical system, explained with reference to FIG. 1, of the detectordevice 17, which together with the front optics 12 and the opticalproperties of the eye 3 defines numerical aperture NA of imaging of theretina 2. The optical system 14 constitutes, thus, a detection opticalsystem 14.

Detection optical system 14 focuses the measurement radiation M into anintermediate image plane, in which a diaphragm 15 is located. Diaphragm15 defines the size of the object field, from which measurementradiation M is collected at the retina 2. Taking into account theimaging scale of detection optical system 14, front optics 12 and eye 3,the size of the diaphragm 15 corresponds exactly to the size of theobject field on the retina 2, from which measurement radiation M iscollected.

The diaphragm 15, as will be explained below, is formed as a multi-holediaphragm which, together with subsequent components, to be explained ineven more detail later, images a plurality of object spots on the retinaonto a corresponding number of detector spots on detector 19. Detectoris designed such that each detector spot, in one direction, is covered 4to 100 pixels, in other examples from 5 to 50 or 5 to 40 pixels. Thedetector thus samples each spot with respect to its intensitydistribution using individual detector areas. The significance of thissampling establishing holographic OCT will be discussed below.

According to FIGS. 3a-d illumination radiation B provided at the end ofthe fiber 9 is distributed onto spots and coupled in over the wholeusable pupil P (diameter 4.5 mm) of the eye 3. In this way, both theupper power limits at the cornea of the eye 3 and the upper power limitsat the retina 2 of the eye are complied with. If the spots are formedsuch that they have a minimum spacing of 2 mm on the retina 2 (ingeneral terms: a spacing which approximately corresponds to thedetectable depth area of the OCT), photons scattered multiple times aresuppressed as effectively as in the case of a confocal OCT, which woulddeflect an individual spot. This solves the problem of scattered light.

Illumination radiation B coupled out of fiber 9 is collimated bycollimator lens 31 and then bundled onto multi-hole diaphragm 34 usingmulti-lens array 32 and field lens 33. Multi-hole diaphragm 34 specifiespattern, spacing and size of the illumination spots on the retina 2, asit lies in a plane which is, because of the subsequent optical system35, 12 a, 12 b, conjugated to the object plane on the retina 2. Opticalsystems are for example, configured such that both the beam splitter 11and the scanner 13 lie close to a pupil of illumination beam path.Optional field lens 33 in front of the multi-hole diaphragm 34 ensuresthat in the plane of pupil P of the eye 3 the radiation is distributeduniformly over the whole pupil P, i.e. over the diameter of 4.5 mm, withthe result that no points exist there in which the maximum radiationintensity might be exceeded.

Measurement radiation M backscattered on the retina 2 is imaged onto thedetector 19 by way of the front optics comprising optical systems 12 a,12 b via the intermediate image plane 26 and the scanner 13 as well asthe beam splitter 11, which both lie close to or in a conjugated pupilplane which is conjugated to the plane of pupil P of the eye 3; ofcourse after reference radiation has been coupled-in by the detectoroptical system 14 (in the section between the dotted double lines). Themulti-hole diaphragm 15 in FIG. 3 realises the diaphragm described withreference to FIG. 1. A downstream multi-lens array 36 images themeasurement radiation M from the individual diaphragm openings of themulti-hole diaphragm 15 onto the detector 19. The multi-lens array 36ensures that the measurement radiation M which is incident through theindividual diaphragm openings of the multi-hole diaphragm 15 no longermixes before it reaches the respectively allocated pixels of thedetector 19. Each hole of the multi-hole diaphragm 36 illuminates agroup of several pixels of the detector 19 with each group forming onedetector area.

FIGS. 3b-d show some options how to couple-in reference radiation R.Reference radiation R originates by way of example from an end of anoptical fiber 70 which forms the end of the additional pathlengththrough which the reference radiation R is sent. A free beam guiding isequally possible for all embodiments.

With reference to FIG. 3b the reference radiation is superimposed withthe measurement radiation to perform a balanced detection. Balanceddetection is also done in the embodiment of FIG. 3c and an off-axisdetection is carried out in the embodiment of FIG. 3 d.

As FIG. 3b shows, the measurement radiation is superimposed, byapplication of a beam splitter 71, with the reference radiation R whichis guided by optical fiber 70 and in advance is distributed, using alens 72 and a multi-lens array 36 c, into a plurality of spots, whereinthe focal plane of the multi-lens array 36 c lies in an intermediateimage plane 26. A multi-hole diaphragm 15 c can optionally lie there.Optical system 14 c collimates the radiation and guides it to beamsuperimposer 71. The latter superimposes the measurement radiation Mwith the reference radiation R and guides the radiation to two detectors19 a, 19 b, upstream of each of which a optical system 14 a, 14 b, amulti-hole diaphragm 15 a, 15 b and a multi-lens array 36 a, 36 b arearranged. The multi-lens arrays 36 a, 36 b in front of the detectors 19a, 19 b and the multi-lens array 36 c are matched to each other and tothe multi-lens array 32 as well as the multi-hole diaphragm 34 of theillumination beam path.

The embodiment represented in FIG. 3b can be used for all multi-spotsystems of the invention which are equipped with a balanced detection.The measurement radiation M is superimposed in the detection beam pathclose to the pupil coherently with the reference radiation R and thenimaged into and detected in two independent beam paths with the twodetectors 19 a, 19 b. The multi-hole diaphragms 15 a, 15 b areconjugated to the multi-hole diaphragm 34 of the illumination beam path.The measurement radiation M which passes through the multi-holediaphragm 15 a or 15 b is collimated by the multi-lens array 36 a or 36b and recorded with the detector 19 a or 19 b.

If the illumination beam path utilizes a pupil in the eye with adiameter of approx. 1.2 mm and the detection beam path uses a pupil witha diameter of 4.5 mm, the microlenses of the multi-lens arrays 36 a-b inthe detection have a focal length that is 4.5/1.2=3.75 times smallerthan the microlenses of the multi-lens array 32. The angular spectrum ofthe radiation at the area detectors 19 a, 19 b of the various spots thenprecisely fills the sensor, without there being an overlap or gaps. Theimaging scale between the image plane of the retina 2 and the multi-holediaphragms 15 a, 15 b of the detection is chosen such that a desirednumber of pixels covers and detects each individual spot which isgenerated by one microlens of multi-lens array 36 a, 36 b, for exampleten pixels per spot are used. The detection is done close to the pupil,i.e. detectors 19 a, 19 b lie in a plane which is conjugated to pupilplane P. The multi-hole diaphragms 15 a, 15 b on the other hand lie inintermediate image plane 26 conjugated to image plane (plane of theretina 2).

To have coherent detection, each bundle of measurement beams issuperimposed with a bundle of reference beams at an identical aperture.This is achieved by collimating reference radiation R, which emergesfrom optical fiber 70, with lens 72 and focusing it by the multi-lensarray 36 c into intermediate image plane 26. A reference wave formsthere in form of a multi-spot pattern which is imaged onto multi-holediaphragms 15 a, 15 b with the aid of the further lens 14 c as well aslenses 14 a, 14 b in the superimposed beam path section of referenceradiation R and measurement radiation M. Lens 14 c for example formswith the lenses 14 a and 14 b a 4f type arrangement.

If each spot illuminates a field with a diameter of approximately 20 μmon the retina 2 and these spots have a spacing of approx. 2 mm, themulti-lens arrays 36 a, 36 b utilize comparatively small effective fieldangles. It is then not necessary for the detectors 19 a, 19 b to bestrictly in the focal planes of the microlenses of the multi-lens arrays36 a, 36 b, rather they can also be at a greater distance. Phasevariances which can may occur over the area detectors can be numericallycompensated for after the coherent detection.

If distance between microlenses of the multi-lens array 36 a, 36 b andthe detector 19 a, 19 b can be larger, a particularly simple detectionarrangement for the balanced detection is possible, as FIG. 3c shows. Itis sufficient then to collimate the reference radiation R with the lens72 only such that the detectors 19 a, 19 b are illuminated in 2D. Only asingle multi-hole diaphragm 15, with one multi-lens array 36, which isupstream of the beam combiner 71 is needed then.

FIG. 3d shows an off-axis detection in which there is sufficient spacebetween the multi-lens array 36 of the detection beam path 17 and thedetector 19 to superimpose with measurement radiation and referenceradiation R. In this case too, if the spacing between multi-lens array36 and detector 19 is larger than the focal length of the microlenses ofthe multi-lens array 36 prescribe, phase variations forming over thedetector 19 can be numerically compensated when processing data.

For the principle of off-axis detection it is generally preferred, forexample, to implement the multi-lens array 36 using anamorphiccylindrical lenses on the front and back side of a plane-parallelsubstrate layer plate of certain thickness. This arrangement, togetherwith a rectangular arrangement of the microlenses in the multi-lensarray 36, also makes it possible to illuminate the camera pixels of thedetector 19 without losses in off-axis, even if more pixels (e.g. 2-3times) are needed in the off-axis direction for imaging with the sameaperture values.

In off-axis detection the angle relative to the optical axis is chosenaccording to various detection parameters. The smaller the angle, thelarger the spacing between multi-lens array 36 and detector 19. Spacingsthat are too large and angles that are too small have the result thatthe phase variances forming can no longer be numerically correctedsufficiently well. An angle that is too large on the other hand has theresult that the coherence of the superimposition may be lost. The use ofa TIRF prism as beam combiner 71 represents a particularly goodcompromise. This prism is constituted by two glass prisms with a smallair gap in-between, which is drawn in schematically in FIG. 3d for thebeam superimposer 71. The glass prisms are dimensioned such that themeasurement radiation M is incident on the air gap at an angle of 45°and can still be transmitted, whereas the reference radiation R, becauseof the angular offset, is incident at more than 45° and totallyreflected, thus reaches the detector 19 in full. The angle of 45° isonly one example. The decisive thing is that, for the measurement beam,the limit angle for total internal reflection is not exceeded, with theresult that the measurement radiation M can pass through the TIRF prism,while the light from the reference is incident on the glass-airinterface at a larger angle than the limit angle, with the result thatthe reference light R is reflected and therefore is guided to thedetector. The limit angle for prisms (with a refractive index ofn(prism)>1) with an air gap (with a refractive index of n(air)=1) istheta.limit=arcsin(n(air)/n(prism)). At the same time the angle of thereference light R on the sensor must, naturally, correspond to the anglefor an off-axis detection. This results from the width of the pixels andthe number of pixels, over which a phase amplitude of 2*pi isdistributed.

As already explained above, the image information is present in a pupilof the beam path in the form of angle information, and the intensitydistribution in the pupil is generally entirely uniform. It is thereforea preferred in an example embodiment to arrange optical elements whichare to act equally on all structures to be imaged in a pupil. Suchelements are, for example, the scanner 13 and the beam splitter 11.However, it is not mandatory to arrange these elements entirely andexclusively in a conjugated pupil plane. In the embodiment of FIG. 3 itis sufficient for example to arrange these elements such that the beamswhich belong to neighbouring holes of the multi-hole diaphragm 34 arealready superimposed. This is the case when the corresponding edge beamsintersect.

Similarly, an embodiment is preferred for example, in which lenses orother elements which can generate reflections, are arranged wherepossible outside a conjugated pupil plane. Here too, this provision isnot to be understood as imperative. It is sufficient to arrange suchelements in areas in which bundles of beams of neighbouring holes of themulti-hole diaphragm 34 do not yet start to overlap, thus their edgebeams have not yet intersected. In the case of the embodiment of FIG. 4the term “close to the pupil” or “far away from the pupil” thus relatesto location along the optical axis at which edge beams of neighbouringholes of the multi-hole diaphragm 34 start/end to intersect.

The scanner 13 of OCT 1 of FIGS. 1 to 3 is located for example, in orclose to a pupil plane of the detection beam path as well as also of theillumination beam path. This pupil plane is conjugated to the plane ofthe pupil P of the eye 3.

The front optics 12 optionally comprises, as shown by way of example forthe embodiment of FIG. 2, opticals 12 a and 12 b which, together, form a4f type optical system. Opticals 12 a is then an ophthalmoscopic lensand opticals 12 b is a scan lens. This 4f type optical system imagespupil P of eye 3 to pupil plane which is conjugated to the plane ofpupil P and in which the scanner 13 lies. It is not necessary to placethe scanner 13 exactly in this conjugated pupil plane, but hasadvantages. An intermediate image plane 26 is located between the planeof pupil P of the eye 3 and pupil plane conjugated thereto. The beamsplitter 11 is located, because of its proximity to the scanner 13,likewise close to conjugated pupil plane. It is also possible to placethe beam splitter 11 in this conjugated pupil plane if the scanner 13 ismoved out of conjugated pupil plane.

In an embodiment the beam splitter 11 is formed by a polarizing beamsplitter. This is then preceded in the imaging direction by a lambda/4plate 27 (cf. FIG. 2). This embodiment will be discussed in thefollowing.

The detector optical system is preferably, for example, likewise formedas a 4f type optical system. It provides a further intermediate imageplane 26 in which the diaphragm 15 lies. The intermediate image plane 26is conjugated to the object plane in which the retina 2 to be imagedlies.

Diaphragm 15, 15 a, 15 b has two functions in all embodiments. Firstlyit suppresses scattered light, whereby the contrast on the detectordevice 17 is improved. The diaphragm acts, in this respect, similarly toa confocal diaphragm for confocally sampling OCT. The detector 19 ispositioned, because of the detector optical system, for example in aplane which is conjugated to the pupil plane of the eye, or close tothis plane. This arrangement is advantageous, but not mandatory. It hasthe advantage that the phase function of the electromagnetic field canbe sampled simply. The maximum spatial frequency in the plane of thedetector 19 is predefined by the object field size on the retina 2 andthus ultimately by the size of the diaphragm 15 in the intermediateimage plane 26. The diaphragm 15 thus ensures, on the other hand, aparticularly favourable signal generation and processing.

In all embodiments of the OCT the detector has, per hole of themulti-hole diaphragm 15, a pixel group of, for example, 4 to 100 pixels,in another example—5 to 50, in a further example 5 to 40 pixels in eachdirection. FIG. 5 shows, in a top view of the detector pixels 43, thatthe arrangement of the pixels 43 need not necessarily be rectangular,but a hexagonal arrangement of the pixels 43 also comes intoconsideration. The pixel pattern can thus be chosen freely.

In the state of the art, holoscopic OCT systems are known which havedetectors with 100 to 4000 pixels per direction. These pixel numbers aredeliberately not used here. The number of pixels is linked to thenecessary illumination brightness, the measurement speed and thesuppression of multiple scattering.

In an example embodiment of the OCT 1 aberrations are corrected. Thepixels of the detector 19 are also referred to as channels in thefollowing. The measurement signal is distributed over these severalchannels. If the detector 19, according to an example embodiment, liesin a conjugated pupil plane, each channel of the detector containsmeasurement radiation M from various angles which was scattered insidethe retina 2. The spatial resolution of the detector 19 makes itpossible to detect the distribution of the measurement radiation in thepupil P for each spot. The following explanation refers to only one ofthese spots. Aberrations affect this distribution. Aberrations caused bythe eye 3 often take on a no longer tolerable dimension if, in the planeof the pupil P of the eye 3, a cross-section larger than 1.5 mm indiameter is covered. Such a larger area would, however, be desirable inrespect of the lateral resolution. Without spatial resolution in theconjugated pupil plane, phase differences would be mixed and averagedout in the then single detection channel when a larger pupil is utilizedon the eye 3.

The corresponding Zernike polynomials which describe these aberrationsare shown in FIG. 4, which shows top views 37 to 42 of the intensitydistribution in a conjugated pupil plane. Further, a grid of thedetector with 5×5 channels (or pixels) per spot is represented. Thepixels sample the pupil P and thus make it possible to distinguish phasedifferences within the pupil P.

The maximally resolvable phase differences depend on the number ofchannels per spot. The inventors found out that the number ofdistinguishable phase differences in this plane is given by the numberof channels per direction multiplied by pi. In the case of five channelsper direction, as represented in FIG. 4, polynomials up to Z₄ ^(m) canbe distinguished per spot, wherein m can take on the values 0 (sphere),±2 and ±4. This applies to channels that are infinitesimally small interms of surface area. In reality, naturally, they have a particularsize. Then, the measurement signal detected in a channel corresponds toan averaging of interference signal over the surface of the respectivechannel (e.g. pixel surface). The theoretically possible, maximum orderof the Zernike polynomial can thus only be achieved if, for each spot,the phase of the signal within a channel varies less than pi. Theinventors found out that, in the case of an average OCT wavelength of1060 nm for the astigmatism caused by the eye the phase differences arerecognizable with uniformly spatially distributed channels if, in thecase of five channels per spot, the condition 2*pi/(5 channels peraberration period)≦pi is met. One period of minima and maxima lieswithin the aperture then. For higher orders, the following is true:0.6*2*pi/(5 channels per period of the aberration)=1.2*pi/(5/1.5)≦pi forthe third order and 0.5*2*pi/(5 channels per period of theaberration)=1.0*pi/(5/2)≦pi for the fourth order.

These considerations show that an area detector with at least fivechannels per direction and spot is capable of resolving at least theastigmatism and the third-order aberrations. A higher number of channelsmakes it possible to detect even higher orders of the aberration.

The above calculations took into consideration only one spatialdirection. As FIG. 4 shows, the aberrations generally have atwo-dimensional pattern. FIG. 4 shows, in top view 37, the first-orderaberration which is also called “piston”, top view 38 shows “tilt”aberration, top view 39 shows “tip” aberration, top view 41 shows“defocus” aberration and top views 40 and 42 show aberrations of“astigmatism” type. As can be seen, most aberrations have atwo-dimensionally distributed pattern, whereby the phase variation isalso two-dimensional. These patterns can be detected and corrected foreach spot by the spatially resolving detector 19, e.g. having at least 5channels/pixels in each direction.

The aberrations bring about, for each detector channel c, a phase θ_(c):U_(sample,c):=U_(sample)*e^(iθ) ^(c) . It results from a thickness δdand a refractive index δn of the material of the eye being passedthrough (e.g. cornea, aqueous humour, lens, vitreous humour), which inreality differs from a theoretical, aberration-free eye:

θ_(c)(k)=δn(k)*k*δd _(c)

The detected signal is thus shifted by the aberration-related phase:

I _(bd,c)(k)=4*u _(s) *u _(r)*cos(k*Δz δn(k)*k*δd _(c))=4*u _(s) u_(r)*cos(k*(Δz−δn(k)δd _(c)))

At monochromatic radiation of 780 nm the eye causes wavefrontaberrations of up to 0.7 μm, which lead to a phase shift of 2*pi (ifdefocus is disregarded). Such phase shift corresponds to a thicknessdeviation between lens and aqueous humour (these are the elements withthe largest refractive index differences in the eye), which of thefollowing value:

${\delta \; d} = {{2{pi}*\frac{780\mspace{11mu} {nm}}{2\; {pi}*\delta \; {n\left( {780\mspace{11mu} {nm}} \right)}}} = {\frac{780\mspace{11mu} {nm}}{{n_{lens}\left( {780\mspace{11mu} {nm}} \right)} - {n_{aqueous}\mspace{11mu} \left( {780\mspace{11mu} {nm}} \right)}} \approx \frac{780\mspace{11mu} {nm}}{1.415 - 1.334} \approx {10\mu \; {m.}}}}$

With known dispersion data, the following results:

$\begin{matrix}{{\theta_{c}\left( {\lambda_{0} = {1060\mspace{11mu} {nm}}} \right)} = {\frac{2\; {pi}}{1060\mspace{11mu} {nm}}*\left( {{n_{lens}\; \left( {1060\mspace{11mu} {nm}} \right)} -} \right.}} \\{\left. {n_{aqueous}\mspace{11mu} \left( {1060\mspace{11mu} {nm}} \right)} \right)*\delta \; d_{c}} \\{= {\frac{2\; {pi}}{1060\mspace{11mu} {nm}}*\left( {1.4104 - 1.3301} \right)*10\mu \; m}} \\{= {1.516\; {pi}}}\end{matrix}$ or θ_(c)(k₀) = k₀ * 0.8034  μ m.

If a wavelength range of Δλ=50 nm is chirped, the phase differences ofthe associated wave numbers (k₀±Δk) are:

${\theta_{c}\left( {k_{0} + {\Delta \; k}} \right)} = {{\frac{2\; {pi}}{1110\mspace{11mu} {nm}}*\left( {{n_{lens}\; \left( {1110\mspace{11mu} {nm}} \right)} - {n_{aqueous}\mspace{11mu} \left( {1110\mspace{11mu} {nm}} \right)}} \right)*\delta \; d_{c}} = {{\frac{2\; {pi}}{1060\mspace{11mu} {nm}}*\left( {1.4099 - 1.3297} \right)*10\; \mu \; m} = {{1.445{pi}} = {\left( {k_{0} + {\Delta \; k}} \right)*0.8022\mspace{11mu} \mu \; m}}}}$     and${\theta_{c}\left( {k_{0} - {\Delta \; k}} \right)} = {{\frac{2\; {pi}}{1010\mspace{11mu} {nm}}*\left( {{n_{lens}\; \left( {1010\mspace{11mu} {nm}} \right)} - {n_{aqueous}\mspace{11mu} \left( {1010\mspace{11mu} {nm}} \right)}} \right)*\delta \; d_{c}} = {{\frac{2\; {pi}}{1060\mspace{11mu} {nm}}*\left( {1.4098 - 1.3305} \right)*10\; \mu \; m} = {{1.594{pi}} = {\left( {k_{0} - {\Delta \; k}} \right)*0.8048\mspace{11mu} \mu \; {m.}}}}}$

These calculations show that, in a sufficiently close approximation, thephase shifts which are caused by the aberrations vary linearly with thewave number k within a wavelength tuning. The detected measurementsignal can thus be written as follows:

I _(bd,c)(k)=4*u _(s) *u _(r)*cos(k*(Δz−δn(k ₀)δd _(c))).

A Fourier transform for the measured wave numbers k give the axialdistribution, i.e. the distribution in the z direction for thescattering tissue. Relative to an aberration-free system the axialdistribution is shifted by the value δn(k₀)δd_(c) for each channel c ofthe detector. FIG. 8 shows a corresponding simulation example, in whichthe z-coordinate is plotted on the x axis and the signal intensity isplotted on the y axis. The curves 55 to 58 correspond to four channels cof the detector. It can be assumed that, in most of tissue, variationsof the axial scattering profile are small within a pupil size of 5 mm ofthe eye 3. Profile differences for the channels c are therefore duemainly to aberrations which shift the profile axially. It is thereforeprovided to relate the channel phases θ_(c)(k₀), caused by aberrations,to a central channel (for example the channel which lies in the centreof the detector and which corresponds to a perpendicular incidence onthe sample). The measured intensities for a frequency determination aremultiplied by the phase factor for correcting the aberrations. The phasefactor is e^(−iθ) ^(c) ^((k) ⁰ ⁾.

Each channel of the detector has a particular position relative to theretina 2. The interference signal can be recorded for each wave numberk=2*pi*n/λ during the wavelength shift/chirp of the laser, wherein n isthe refractive index of the medium and λ is the wavelength. As known toa person skilled in the art of conventional OCT systems, the measurementsignals are Fourier-transformed in respect of the wave numbers, and thedepth distribution of the scattering layers is calculated. Therelationship Δφ=k*Δz is used, wherein Δz is the distance of a scatteringlayer from a reference layer from which the measurement radiation wastransmitted to the detector along a pathlength which is identical to thepathlength of the reference radiation beam path.

Because of the lateral extent of the detector 19 per spot, however, theoptical pathlength is not identical for all pixels of a spot, as FIG. 6shows. It can be seen there that for each spot in the detector 19, forwhich five pixels or channels 46, 47, 48, 49 and 50 drawn in by way ofexample, differ in respect of the optical pathlength from a particularpoint in the tissue 44 sampled by the spot. The wavefronts for centralchannel 48 are drawn in with continuous lines. They are perpendicular tothe optical axis of the beam ray and are between the observed point,which is drawn in at the very bottom in the structure 44, and thechannel 48. For this central channel 48 the beam ray runs along theoptical axis. For a channel placed further out, for example outerchannel 50, the main beam ray runs at an angle α relative to the opticalaxis, with the result that the path length has value d for the centralchannel 48 and value d*cos(α) for outer channel 50. For outer channel50, FIG. 6 shows the wavefronts and main beams in dotted line. Further,FIG. 6 shows the eye lens 45 by way of example. The depth is referred tothe main plane of the eye lens, as the jump in the refractive indexthereof gives a reference point for the measurement. As FIG. 6 clearlyshows, pixels/channels which lie further out collect radiation whichcovered a longer path through the medium. This has some effect onreconstruction of the image information as FIG. 7 shows by way ofexample. Signal curves 51 to 54 are shown there for four channels of aspot. The plotting corresponds to that of FIG. 8, i.e. the depthcoordinate z is plotted on the x axis, the intensity is plotted on the yaxis. As can be seen, the individual curves are not only shifted in thez direction, they are also compressed for pixels lying further out. Thecurve 54 is the measurement signal of the central pixel 48, and thecurves 53, 52 and 51 are measurement signals from channels lyingfurther, respectively.

A measurement error caused by this effect is corrected in an exampleembodiment, in order to obtain a particularly good image. The geometriceffect is for example corrected by a rescaling from z to z*cos(α_(c))for each spot, wherein α_(c) is the angle which the c^(th) channel hasrelative to the optical axis. The angle α is measured against a virtualposition of the detector 19 in which virtual position the detector isplaced directly in front of the eye, of course, while taking intoaccount the imaging scale. In the case of a detector which lies exactlyin a plane conjugated to the pupil plane of the eye, in this way thevirtual position of the detector is exactly in the plane of the pupil Pof the eye 3 with dimensions of the detector modified according to theimaging scale.

During aberration correction of the reconstruction each channel isreconstructed independently. A cross-correlation is calculated in axialdirection, i.e. in depth direction, in order to determine the relativephase offset between the individual channels. A reconstruction of thelateral image for each channel (optionally, as will be described below,taking into account the scanning process) and then of the phase gradientsupplies a lateral offset in the image which is obtained for a givenposition of the scanner. This image is also called the pupil channelpartial image in the following. In an embodiment the aberration isdetermined by application of a lateral cross-correlation of the pupilchannel partial image and in this way the whole aberration phasedistribution is determined and numerically corrected.

The quality of these approaches depends on the sample structure. In thecase of the human eye, a very prominent axial layer structure is found.Laterally relative thereto the structures are relatively coarse, forexample due to blood vessels or the papilla, combined with very finestructures, such as photoreceptors, wherein hardly any structure, withrespect to size and course, lies in-between. In an example embodiment adepth correlation correction is first carried out by using the axiallayer structure in order to correct the majority of the pupil phaseaberrations. Optionally a lateral correlation correction follows, whichutilizes lateral structures, such as for example photoreceptors, whichbecame visible because of the first correction.

The aberrations of the eye are different at different sites on theretina. In principle it is possible to calculate the phase changescaused by aberrations in each channel for all points in a lateral image.In a simplified embodiment it is assumed that aberrations do not varyvery strongly in lateral direction, and aberrations are only calculatedfor few lateral locations on the retina and interpolated forintermediate locations.

If a comparatively large wavelength range is tuned/chirped, it ispreferred, for example, to take into account the dispersion ofaberrations. In this embodiment it is not assumed that the phase shiftschange linearly with the wave number k. A peak in profiles whichoriginates in the OCT image from the retina 2 at the fundus of the eye 3is therefore used in order to compensate for the shift of profilesrelative to each other. Thus, for example, a structure (in the form of apeak) is sought in the curves 51 to 54 of FIG. 7, and the curves arecorrected relative to each other using such reference structure. In thisway aberrations θ_(c)(k₀) can be determined and corrected as describedabove. Alternatively a complex correlation algorithm which is applied toprofiles of the different channels is also possible. In addition to ashift, a scaling (compression or expansion) of the measurement signalscan also be corrected.

Each detecting status position of the scanner 13 gives a partial imageof the retina, the size of which image is predefined by the diaphragm 15(extent and hole size and number) and the front optics 12 and detectoroptical system 14 that cooperate during the imaging of the measurementlight. A Fourier transform of the signal of the channels gives the imageof the sample, but for each spot only in that part which corresponds tothe size of the detected spot in the pupil. In order to generate alarger image, the scanner 13 is provided and operated, which shifts theposition of the imaged object field, thus the object spots on the retina2. The image area of each spot corresponds to a partial image 59 whichhas a centre 60. For a current deflection by the scanner 13 it issufficient, for simplification, to refer to the centre 60 of the partialimage 59. Scanning of multi-spot images is known to persons skilled inthe art, for example in confocal microscopy. However, they are to besupplemented here to the effect that not only lateral information byadjustment of the scanner, but also depth information by tuning of thewavelength of the radiation source is obtained.

This opens different scanning approaches. The scanner can rest duringthe tuning of the wavelength of the light source 4. Before a renewedtuning takes place, the scanner is moved to a new position of the spotpattern, suitably spaced to the previous position. In this way thepositions of the spot pattern can acquire a larger total image 61 of theretina. This approach is shown in FIG. 9 for one depth plane. As aresult individual partial images 59 (each corresponding to one positionof the spot pattern) are combined to the total image 61. The informationof the individual depth planes then result in a three-dimensional imageof a cuboid-shaped region of the retina 2. This is shown by FIG. 10, inwhich three planes 62, 63 and 64 are shown by way of example. Thepartial images 59, which are related to each other in the representationof FIG. 10 by a dot-dashed double arrow, in each case originate from awavelength tune/chirp of the light source 4. As the scanner 13 isstationary during each wavelength tune/chirp and only adjustedin-between, the partial images 59 generated from one wavelengthtune/chirp lie all with their centres 60 of the planes 62 to 64 exactlyone over another.

For particular example embodiments it is preferred to scan continuously,i.e. to adjust the scanner 13 while the wavelength is tuned/chirped.This approach requires synchronization of scanner 13 and wavelengthtune/chip of the light source 4. It is preferred for example, to set thelateral adjustment speed of the scanner 13 such that one wavelength tunecovers at most one partial image 59 in one direction, preferably forexample, not even a full image. Partial image 59 then differs frompartial image of FIG. 9. In this approach the travel length of thescanner 13 can for example correspond to the spacing of the objectspots, with the result that a partial image 59 corresponds exactly tothe shift of the spot pattern within one period, i.e. the spacingbetween two neighbouring individual object spots. Thus, for planes 62,63 and 64 the position of their centres 60 changes, as the partialimages 59 in the planes are based on the different wavelengths in theFourier transform. Unlike in the embodiment of FIG. 10 a provisionaltotal image 61 is obtained which is not a rectangular cuboid, but e.g. ano longer rectangular parallelepiped, because of the adjustment of thescanner 13 during the wavelength tuning. An example embodiment correctsthis effect for the imaging by trimming the parallelepiped down to acuboid.

There are various possibilities for taking into account the simultaneityof wavelength tuning and lateral shift. If the detector lies close to anintermediate image plane, thus in a plane conjugated to the retina, thedata of the three-dimensional parallelepiped are shifted relative toeach other. For each wave number k_(i) an image of the sample can beassembled, wherein I_(i)=I(k_(i), x, y) applies. These images I_(i) areoffset a little relative to each other. As the allocation betweenlateral scanning position and wave number is known, the wavelengthtuning can be assembled correspondingly for each location (x, y) in thesample. In this way the three-dimensional data are simply assembled.

In embodiments in which the detector is located in or close to theconjugated pupil plane, it measures the Fourier transform of theintensity distribution in the object plane (retina 2). A shift in theobject plane leads to a phase ramp in the detector plane. The correctionof the simultaneous lateral adjustment by the scanner 13 and thewavelength tuning by the light source 4 is therefore obtained from amultiplication of the detector signal by a time-dependent phase rampwhich is proportional to the scanning speed and the spacing betweenpupil partial channel and optical axis in the pupil plane.

The embodiments of FIGS. 1 to 3 have the result that the illuminationand the receipt of the measurement light are no longer coupled to eachother with respect to the optical properties and in particular the pupilsize. In this way it is possible to fine-tune the illumination. Forexample, a Bessel-type illumination can be combined with a top-hatcross-sectional profile for the detection. This provides embodimentscombining a large illumination depth, i.e. an illumination which isunchanged over a large z area, with high numerical aperture of theimaging. In case of an identical numerical aperture, for example with aGaussian beam, an illumination focus with an extent of 1 mm in zdirection would be achieved. In the case of a Bessel-type illuminationan extent of 2 to 3 mm in the z direction is obtained. In this way theoptical resolution can be increased by 10 to 30% if the detection isdone with a top-hat profile.

In a further embodiment of the OCT the beam splitter 11 effectspolarization splitting. This is usually considered to be disadvantageousin the state of the art, and an intensity splitting is preferred.However polarization splitting is surprisingly advantageous for the OCTof the present invention, as polarized radiation entering the eye ischanged therein with respect to its polarization state. Differentstructures of the eye have a different polarization changing effect,with the result that the polarization state of the backscattered signalis not unambiguously or clearly defined, but consists of components withdifferent polarization states. This consideration led the state of theart to carrying out an intensity splitting, simply because thebackscattered radiation does not have a clear, defined polarizationstate. However, the inventors found out that only beam constituentswhich have the same polarization state can interfere with each otherwhen the measurement light is superimposed with the reference light. Itis the polarization state, of the reference light which predefines whatportion of the measurement light can be utilized. Non-interferingportions form background noise on the detector.

The polarization splitting is now explained with reference to theembodiment of FIG. 2, but is not limited to the features otherwiserealized there and can, for example, also be used in the embodimentaccording to FIG. 1. The illumination radiation B is linearly polarizedafter the polarization splitter 11. The lambda/4 plate 27, as shown inFIG. 2, generates circularly polarized illumination radiation B fallingonto the eye 3. Backscattered measurement radiation M, which is likewisecircularly polarized, is again linearly polarized by the lambda/4 plate,wherein the polarization direction is rotated through 90 degreesrelative to the polarization direction which the illumination radiationB has which was emitted by the polarization splitter 11. The measurementradiation M thus passes through the polarization splitter 11 withoutdeflection and interferes with the reference radiation R if the latterhas the same polarization. This is the case when reference radiation Rand illumination radiation B are identically linearly polarized afterthe dividing the source radiation. This is also the case if referenceradiation R and illumination radiation B are circularly polarized afterthe dividing the source radiation and the reference radiation islinearly polarized identically to the measurement radiation M before thesuperimposition. Finally, it is important that the polarizationsplitting (e.g. by polarization splitter 11 and plate 27) conditions themeasurement radiation M and the reference beam path conditions thereference radiation R identical to such extent that both radiations havethe same polarization state on the detector.

This increases the signal-to-noise ratio, as only those parts of themeasurement light that are capable of interfering with the referencelight are forwarded by the beam splitter 11 to the detector device 17.Finally, the polarization splitting and rejection of a part of themeasurement radiation M at the beam splitter 11, which are bothdisadvantageous at first glance, increase the signal quality.

In a further embodiment the OCT uses the fact that the illuminationoptical system 10 can to place the focus of the illumination radiation Bat another z position than the focus which is predefined by the detectoroptical system 14 for the collection of the measurement radiation M.Because of multiple scatterings in the retina, measurement radiation Mfrom the retina can have a pathlength suitable for interference, but canpropagate in another direction, which would limit the lateral resolutionin terms of the depth. This effect can be compensated for by differentfocal depth planes for illumination and detection. The depth resolutionis optimized.

For image reconstruction from the detector signals the currentwavelength must be known according to the FD-OCT principle. Thiswavelength or the corresponding wavenumber k can be derived from controlof the light source 4. Alternatively it is possible to couple out a beamportion and detect its wavelength, in order to better know the currentwavelength or the status of the wavelength chirp.

Perpendicularly to the sampling direction, detector channels can becombined in order to reduce speckles. This is particularly advantageousif only z-sections through the retina are desired.

For a coarsely resolved image, e.g. for a preview image, it is possibleto combine all or several detector channels for each spot. This can bedone after the corrections (e.g. aberration, z-position, total imagegeneration). Resolution of conventional OCT systems is then obtained,however, with a higher signal-to-noise ratio and improved specklebehaviour, simply because the combination is done after one or more ofthe corrections and thus goes beyond a normal pixel binning.

If a detector is used which is only spatially resolving in onedirection, aberrations can be corrected in this direction only. This maybe sufficient for particular applications.

In an embodiment an iris camera is provided which assists the operatorto adjust the device at the eye.

For all embodiments of the described optical coherence tomographs ormethods for optical coherence tomography, the following exampledevelopments can be advantageously used:

Phase errors which form if detector 19, 19 a, 19 b is not locatedexactly in the focal plane of the microlenses of the multi-lens array36, 36 a, 36 b can be corrected numerically.

The microlenses of the multi-lens array and thus ultimately theillumination spots on the retina 2 can be arranged in a square grid orin a hexagonal grid. As, for the multi-hole diaphragms, round openingsare preferred, for example, and the pupil or detection aperture as arule is approximately round, a hexagonal grid enables a further savingof detection pixels, i.e. allows to utilize detectors with fewer pixels.

It is preferred, in an example embodiment, to have, independently of thegrid of the illumination spots on the retina 2, one pixel of the areadetector 19, 19 a, 19 b precisely in the centre of each imaged spot. Inthe case of a hexagonal grid of the illumination spots in combinationwith a rectangular grid of the pixels of detector 19, 19 a, 19 b,therefore, the size of the holes of the multi-hole diaphragm 34 and thusalso of the multi-hole diaphragms 15, 15 a, 15 b should be matched tothe pixel size and the resolution of detector 19, 19 a, 19 b such thatthis condition is met sufficiently, i.e. at least approximately, e.g. to+/−10% of the spot diameter.

Where method steps and/or signal corrections were described above, theseare carried out in the OCT 1 by the control device C which is connectedto the detector, reads its measurement signals and obtains further dataabout the operation of the scanner 13 and the wavelength tuning and/orcontrols these components correspondingly.

1-36. (canceled)
 37. An optical coherence tomograph for examining ascattering sample to be placed in an object area, the optical coherencetomograph comprising: an illumination source that emits source radiationof sweepable wavelength; a dividing element that divides the sourceradiation into a reference beam path and an illumination beam path forilluminating the object area with illuminating radiation; optics in theillumination beam path that distribute the illumination radiation intoseveral object spots and that project these object spots to the objectarea, and a scanner that adjusts the lateral position of the objectspots in the object area; a detection beam path collecting radiationscattered at the object spots as measurement radiation and superimposingthis measurement radiation with reference radiation guided through thereference beam path and guiding the superimposed radiations to aspatially resolving detector comprising pixels, wherein the measurementradiation from an individual object spot is guided to a detector spotcovering several pixels of the detector with the detector generatingsignals therefrom; and a control device processing the signals generatedby the detector and generating therefrom an image of a sample providedin the object area.
 38. The optical coherence tomograph according toclaim 37, wherein the optics comprise a multi-lens array thatdistributes the illumination a into the several object spots.
 39. Theoptical coherence tomograph according to claim 37, wherein the opticscomprise a first multi-hole diaphragm that distributes the illuminationradiation into the several object spots.
 40. The optical coherencetomograph according to claim 37, wherein the detection beam pathcomprises a second multi-hole diaphragm defining an object spot fieldsize for the individual objects spots in the object area, from whichobject spot fields measurement radiation reaches the detector.
 41. Theoptical coherence tomograph according to claim 40, wherein the secondmulti-hole diaphragm is located in or close to an intermediate imageplane of the collecting of measurement radiation.
 42. The opticalcoherence tomograph according to claim 37, wherein a multi-lens array islocated upstream of the detector, which multi-lens array bundlesradiation from each object spot to the assigned detector spot.
 43. Theoptical coherence tomograph according to claim 37, wherein the detectoris located in an image plane of the collecting of the measurementradiation.
 44. The optical coherence tomograph according to claim 37,wherein the detector is located in a far field of the object area. 45.The optical coherence tomograph according to claim 37, wherein thedetection beam path comprises a beam splitter that splits theillumination radiation from the measurement radiation scattered in theobject area.
 46. The optical coherence tomograph according to claim 45,wherein the beam splitter comprises a polarizing beam splitter and alambda/4 plate is located between the object area and the beam splitter,the lambda/4 plate cooperating with the polarizing beam splitter tofilter the measurement radiation regarding a polarization state of themeasurement radiation.
 47. The optical coherence tomograph according toclaim 37, wherein the illumination beam path comprises an opticalelement which acts on the illumination radiation only and which definesthe numerical aperture of illuminating the object area independentlyfrom the numerical aperture of collecting the measurement radiation. 48.The optical coherence tomograph according to claim 37, wherein thereference beam path comprises a multi-lens array which bundles thereference radiation into a multi-spot pattern adapted to the objectspots at which the measurement radiation is collected.
 49. The opticalcoherence tomograph according to claim 37, wherein the scanner shiftsthe lateral position of the spots during a wavelength sweep of thesource radiation and generates a scan signal indicating a deflectionstatus of the scanner, wherein the control device is connected to thescanner and to the radiation source and reads out a wavelength signalindicating an actual wavelength of the source radiation and consequentlyof the illumination radiation and connected to the detector and readsout measurement signals for each pixel, wherein the control devicegenerates from the wavelength signal and the measurement signals partialimages of a sample located in the object area and evaluates the scansignal to combine the partial images into a 3D total image.
 50. Theoptical coherence tomograph according to claim 37, wherein the sampleexamined comprises a human eye.
 51. The optical coherence tomographaccording to claim 50, wherein the sample examined comprises a retina ofthe human eye.
 52. The optical coherence tomograph according to claim50, wherein the illumination radiation is uniformly distributed over across section covered in the pupil of the eye, however, allowing forintensity fluctuations of +/−10%.
 53. The optical coherence tomographaccording to claim 37, wherein, the optical coherence tomograph isstructured to examine the retina; the illumination beam path comprises alight-distributing element that distributes the illumination radiationover the spots and that illuminates the retina, the illumination andmeasurement beam path comprise a shared front optic and a shared beamsplitter that split off the measurement radiation collected at the eyefrom the illumination radiation guided to the eye, wherein the beamsplitter guides the split off measurement radiation to the detectionbeam path, the reference beam path provides an optical path length forthe reference radiation, which optical path length corresponds to anoptical distance from the splitting element to the object spots and backto a point of superposition, and the detection beam path superimposesthe measurement radiation and the reference radiation at the point ofsuperposition and comprises an optical element acting only on themeasurement radiation and co-operating with the front optics to definethe numerical aperture with which measurement radiation is collectedfrom the eye, wherein further a diaphragm is located upstream of thedetector and in or close to an intermediate image plane and defines thesize of an object field from which measurement radiation reaches thedetector, wherein the diaphragm upstream of the detector is formed as afirst multi-hole diaphragm and a first multi-lens array is locatedbetween the first multi-lens diaphragm and the detector, the firstmulti-lens array bundles radiation emerging from each hole of the firstmulti-diaphragm onto a dedicated pixel area of the detector whichpresents a spatial resolution of 4 to 100 pixels in one direction. 54.The optical coherence tomograph according to claim 53, wherein thededicated pixel area presents a spatial resolution of 2D pixel area of 5to 50 pixels or of 5 to 40 pixels per direction.
 55. A method foroptical coherence tomography for examining a sample, including an eye,wherein the method comprises: providing source radiation, sweeping thewavelength thereof and dividing the source radiation into illuminationradiation and reference radiation; illuminating the sample with theillumination radiation at a multitude of object spots, wherein a scanneris used for shifting the lateral position of the object spots; guidingthe reference radiation through a reference beam path; collectingillumination radiation backscattered in or at the sample in a form ofmeasurement radiation by imaging the object spots to detector spots;superimposing the measurement radiation with reference radiation guidedthrough the reference beam path, wherein the detector spots are fed withsuperimposed radiation; and the detecting an intensity distribution ofthe detector spots.
 56. The method according to claim 55, furthercomprising performing imaging of the detector spots in parallel byfiltering the measurement radiation with a multi-hole diaphragm.
 57. Themethod according to claim 55, further comprising detecting of theintensity distribution by oversampling in which a resolution of theintensity distribution of each detector spot assigned to an object spotis larger than a resolution of illuminating of the sample with theillumination radiation at the multitude of object spots.
 58. The methodaccording to claim 57, further comprising obtaining an image correctionfrom the intensity distribution.
 59. The method according to claim 55,further comprising generating the multitude of object spots by using amulti-lens array.
 60. The method according to claim 55, furthercomprising generating the multitude of object spots by using a firstmulti-hole diaphragm.
 61. The method according to claim 55, furthercomprising imaging the objects spots to the detector spots by using asecond multi-lens diaphragm defining a size of object spot fields of theindividual detector spots in the sample, and detecting the measurementradiation from the object spot fields.
 62. The method according to claim55, further comprising locating the second multi-hole diaphragm close orin an intermediate image plane of the imaging.
 63. The method accordingto claim 55, further comprising providing a multi-lens array upstream ofa detector comprising pixels, wherein the multi-lens array bundles eachdetector spot to several pixels of the detector.
 64. The methodaccording to claim 55, further comprising detecting the intensitydistribution of the detector spots in an image plane of the imaging ofthe object spots.
 65. The method according to claim 55, furthercomprising detecting the intensity distribution of the detector spots ina far field of the sample.
 66. The method according to claim 55, furthercomprising separating the illumination radiation from the measurementradiation scattered in the sample by using a beam splitter.
 67. Themethod according to claim 66, further comprising making or selecting thebeam splitter to be a polarizing beam splitter and further comprisinglocating a lambda/4 plate between the object field and the beam splitterand wherein the lambda/4 plate co-operates with the polarizing beamsplitter to filter the measurement radiation regarding a polarization ofthe measurement radiation.
 68. The method according to claim 55, furthercomprising independently setting the numerical aperture of the imagingof the object field separate from the numerical aperture of thecollecting of the measurement radiation.
 69. The method according toclaim 55, further comprising bundling the reference radiation to amulti-spot pattern which is adapted to the object spots at which themeasurement radiation is collected.
 70. The method according to claim55, further comprising shifting the lateral position of the spots duringwavelength sweeps of the source radiation and generating partial imagesof the object based on the wavelength of the source radiation andconsequently of the illuminating radiation and based on the detectedintensity distribution, and composing the partial images to a 3D totalimage under consideration of the shift of the lateral position of thespots.
 71. The method according to claim 55, further comprisingselecting the sample examined to be a human eye.
 72. The methodaccording to claim 71, further comprising selecting the sample examinedto be a retina.
 73. The method according claim 71, further comprisinguniformly distributing the illumination radiation over the pupil of theeye within a cross section covered by the illumination radiation,however, allowing for intensity fluctuations of not more than +/−10%.